The following background includes information that may be useful in understanding the present subject matter. It is not an admission that any of the information provided herein is prior art or relevant to the presently claimed subject matter, or that any publication specifically or implicitly referenced is prior art.
Magnetic Resonance Imaging (MRI) employs a strong magnetic field that is used to polarize the spin magnetization in a patient's body. The spin magnetization that is most often used in MRI arises from the nuclei of hydrogen atoms within the body. Although the highest concentration of hydrogen atoms within the body is found in water molecules, other compounds found in the body (e.g. lipids, glucose, etc.) are present in sufficient concentration to provide a detectable MR spin magnetization.
When the hydrogen atoms of a patient's body are introduced into the polarizing magnetic field, the spin magnetization of the hydrogen atom nuclei align in one of two states: with the magnetic field, or against the magnetic field. These two states occupy slightly different energy levels in a quantum mechanical system. By convention, the lowest energy level is called the ground state. It should be noted that the population of nuclear spins in the ground state is slightly higher than that of the higher energy state, resulting in a net magnetization of the macroscopic group of nuclei.
The energy difference between the two energy levels is directly proportional to the strength of the polarizing magnetic field. Thus, as the strength of the magnetic field is increased, the energy difference between the two states increases. The energy differences associated with typical MRI systems correspond to electromagnetic waves in the radiofrequency range. The specific frequency associated with the difference is called the Larmor frequency (typically given in MHz). The constant of proportionality that defines the relationship between the polarizing field (typically given in Tesla) and the Larmor frequency is a natural constant called the gyromagnetic ratio. This constant is unique for each MR active element. For Magnetic Resonance Imaging systems used in medicine, polarizing magnetic field fields are typically between 0.5 and 3.0 Tesla. For hydrogen atoms, these polarizing magnetic field strengths result in Larmor frequencies between 21.3 and 127.8 MHz.
If the nuclear spin system immersed in a polarizing magnetic field is subjected to a rotating magnetic field at the Larmor frequency, the spin system will absorb energy and the distribution of nuclear spins in the two energy states will be disturbed. The duration of the rotating magnetic field used to change the distribution of nuclear spins in the two energy states is typically limited, and applied with a strength sufficient to nutate the net spin magnetization from the longitudinal axis (i.e. parallel with the applied polarizing magnetic field) to the transverse plane (i.e. perpendicular to the applied polarizing field). The term “RF pulse” is conventionally used to describe this process since nutation is accomplished with a rotating magnetic field in the radiofrequency range and having a finite duration.
With time, the energy will be emitted by the spin system in a fashion that can be detected with a sensitive pickup coil. This phenomenon is typically called “resonance”. The absorption and re-emission of an RF signal is key to the formation of an MR image.
When an MR signal is created, the frequency of the signal is precisely proportional to the strength of the magnetic field experienced by the nuclear spins. If all of the spins in a patient's body are in an identical magnetic field, then all the spins will resonate with the same frequency. Even though signals come from many different portions of the body, the MR imaging system has no way to distinguish one signal from another.
In order to provide spatial encoding of the MR signals (and hence enable the formation of an image), it is useful to create a transient inhomogeneity in the magnetic field. In typical MRI imaging systems this is accomplished with magnetic field gradient coils. Gradient coils typically are designed to create a magnetic field whose strength varies in a linear fashion over a selected volume within the magnet. Gradient coil sets are typically constructed to permit gradient fields to be created in three orthogonal directions within the bore of the magnet. Typical gradient coils driven by typical gradient amplifiers can generate a magnetic field gradient of 20 mT/m in less than 1 ms, and maintain that gradient with high fidelity for an extended period limited only by the heat dissipation of the gradient coils and amplifier.
A typical imaging system creates an image by employing a sequence of RF and magnetic field gradient pulses to establish a detectable MR signal in a selected plane. This signal is then spatially encoded using magnetic field gradient pulses to impart phase and frequency shifts to the MR signal which reveal the location of the signal source within the bore of the magnet. By selecting pulse sequence repetition times (TR), echo times (TE) and other pulse sequence parameters, the operator can tune the pulse sequence to be sensitive to a variety of intrinsic MR parameters found in the tissue of the patient (e.g. longitudinal relaxation time, T1, Transverse relaxation time, T2, and the like). Many pulse sequences are known to those skilled in the state of the art. These pulse sequences can collect data in two or three dimensions. They can also collect data in Cartesian, radial or spiral frameworks.
One aspect common to all MR imaging pulse sequences is that they employ transient magnetic field gradients. These transient gradient pulses are created by running electrical current through the gradient coils that are located within the bore of the magnet. Current running through these coils creates a mechanical force that results in a small physical displacement of the coil and its structure. Because of the temporal duration of these gradient pulses, acoustic noise is created. Despite aggressive engineering measures to minimize the amplitude of these physical displacements, MR imaging systems can be loud and hearing protection for the patient is required.
Because the amplitude of the MR signal detected by the MR scanner receiver system is weak, MR magnets are typically placed inside a screen room which acts as a Faraday cage. All electrical signals into and out of the screen room go through filters located in a penetration panel to prevent stray RF signals entering into the screen room from the outside. In the absence of a Faraday cage, RF interference manifests itself in an MR image as an elevated noise floor and/or “zipper” artifacts. In extreme cases, RF interference can saturate the preamplifiers of the MR imaging system and prevent the acquisition of MR signals. When the door to the screen room is closed, the magnet is surrounded by the conducting surfaces of a Faraday cage and external radiofrequency sources such as two-way radios, computers, infusion pumps and the like cannot interfere with MR image acquisition.
It is often desirable to perform MR imaging in patients who are attached to accessory equipment such as infusion pumps, respirators, and/or anesthesia equipment. In these circumstances it is often not practical to place these pieces of equipment outside the screen room since direct attachment to the patient is necessary. Unfortunately, these types of accessory equipment often create RF noise which can degrade the MR images acquired from the patient.
In view of the foregoing, it may be understood that the reduction of radiofrequency interference during MR scanning is desirable, and may serve to increase the quality of images obtained with Magnetic Resonance. Furthermore, it is desirable to reduce RF interference without modification of the equipment that creates the RF noise.